Digital x-ray imagers are rapidly replacing x-ray film based detectors in medical imaging and other applications. In a digital x-ray imager, the x-ray signal is converted into either light photons or charge carriers, which are then collected and converted to a measurable electrical signal and digitized by an electronic circuit. The digitized signal is then represented as a discrete grayscale level in an image pixel. A matrix of such grayscale pixels forms an x-ray image. For imagers in which x-rays are first converted into light, a scintillator material is used. The scintillator material generates optical light photons when x-rays are stopped by and interact with the scintillator material. Photosensitive elements such as photodiodes collect light from the scintillator. The total amount of light collected by a photosensitive element affects the image signal and noise levels. The spread of the light determines the optical blur in the image. This, together with the detector pixel size determines the image resolution and is measured in terms of modulation transfer function (MTF).
In conventional imagers for low energy (KV) x-ray imaging, the scintillators are grown as columnar needles of about 10 μm diameter that act as light pipes or guides. However, there is still significant cross talk between neighboring scintillator needles. For higher energy (MV) imaging, significantly thicker scintillators are needed to stop the x-rays. Growing very thick structured needles is difficult and its effectiveness to channel light reduces as cross talk increases.
Conventional MV x-ray imaging or electronic portal imaging devices (EPID) suffer from low detective quantum efficiency (DQE). DQE is a measure of the fidelity of an imaging device in capturing and transferring image information. The range of DQE is 0<DQE<1, where the value of 1 implies that all the image information in the incoming X-rays is captured and no noise is added. Portal imagers are typically used at high energies such as MV and hence require thick scintillators to effectively absorb x-rays. The absorbed x-rays then generate optical photons at the location of their interaction. The optical photons that are generated travel in all directions and can be further reflected and/or refracted from interfaces. The light photons that can reach the photodiodes form the final image. With increased thickness of the scintillator, the light photons undergo extensive spreading and result in image blurring.
To overcome the above problems in MV x-ray imaging, pixellated scintillators have been used. Pixellated scintillators are formed by slicing the scintillator crystal into parallelepipeds, which are then coated with a layer of reflective or absorptive coating. The coated parallelepipeds are then joined back together, with the reflective or absorptive layer sandwiched between adjacent pieces. A pixellated scintillator may limit the spread of light but manufacturing such a matrix is labor intensive and the cost for large area pixellated scintillators is prohibitive. Further, the partition between neighboring pixels may cause light to bounce back into the scintillator pixel and as such, the original directionality of the light ray is lost.
Accordingly, there is a general need for improving the resolution of imaging devices for both KV and MV x-ray imaging. There is a need for an x-ray imaging device and method that can reduce the effect of light spreading in scintillators particularly in configurations where conventional approach results in heavy light losses and high costs.